The present invention relates generally to multi-slice computed tomography (CT) imaging systems, and more particularly, to an apparatus and method of generating x-rays within an imaging tube.
There is a continuous effort to increase computed tomography (CT) imaging system scanning capabilities. This is especially true in CT imaging systems. Customers desire the ability to perform longer scans at high power levels. The increase in scan time at high power levels allows physicians to gather CT images and constructions in a matter of seconds rather than several minutes as with previous CT imaging systems. Although the increase in imaging speed provides improved imaging capability, it causes new constraints and requirements for the functionality of the CT imaging systems.
Referring now to FIG. 1, a cross-sectional view of a traditional CT tube assembly 10 is shown. CT imaging systems include a gantry that rotates at various speeds in order to create a 360xc2x0 image. The gantry contains the CT tube assembly 10, which composes a large portion of the rotating gantry mass. The CT tube assembly 10 generates x-rays across a vacuum gap 12 between a cathode 14 and an anode 16. In order to generate the x-rays, a large voltage potential is created across the vacuum gap 12 allowing electrons, in the form of an electron beam, to be emitted from the cathode 14 to a target 18 of the anode 16. In releasing of the electrons, a filament contained within the cathode 14 is heated to incandescence by passing an electric current therein. The electrons are accelerated by the high voltage potential and impinge on the target 18, whereby they are abruptly slowed down, directed at an impingement angle xcex1 of approximately 90xc2x0, to emit x-rays through CT tube window 19. The high voltage potential produces a large amount of thermal energy not only across the vacuum gap 12 but also in the anode 14.
The anode 14, as with other traditional style CT tube anodes, uses a store-now/dissipate-later approach to thermal management. In order to accommodate this approach the anode 14 is required to have a large mass and a large diameter target. The electron beam impacts the target 18, near a rim 20, essentially normal to the target face 22. The target 18 is rotated about a center axis 24 at approximately 180 Hz or 10,000 rpm to distribute load of the electron beam around a track region 26 of the target 18. Thermal energy generated in the track region 26 is transferred through the target 18 to a thermal storage material, such as graphite, which brazed to a back surface of the target 18. As the anode 14 rotates, thermal energy stored on the back surface of the target 18 dissipates during each revolution of the anode 14, thereby cooling the anode 14.
Traditionally, in order to increase performance of a CT imaging system, thereby increasing the amount and frequency of electron emission for a given duration of time, the diameter and mass of the target is increased. By increasing the diameter and mass of the target, thermal energy storage and radiating surface area of the target is increased for increased cooling.
Increasing the diameter and rotational speeds of the target is limited due to size, mass, and material strength of the target. The stated limitations in combination with a large amount of rotationally induced stress in the target, from instantaneous power being applied over very short durations on the target, also limit linear velocity of the track. Size of the target is also further limited by space constraints in a CT imaging system. An example of a space constraint, is the desire for good angulation capability, in that in cardiac or similar applications the CT system needs to be mobile and position flexible. Other space constraints exist and are commonly known in the art.
Additionally, faster scanning increases the mechanical loads on an entire CT tube, especially anode bearings, thus degrading CT tube component performance. Hence, in order to minimize mechanical loads the ability to increase the mass of the target is limited, which conflicts with the thermal performance of the X-ray tube. Faster scanning in increasing anode surfaces can cause subcooled nucleate boiling further decreasing scanning quality.
There is a continuous desire to perform CT scans at increased rates, thus requiring more instantaneous power to be applied on the target over very short durations potentially causing increased thermal energy. It would therefore be desirable to provide an apparatus and method of generating x-rays within an x-ray tube that provides increased scanning speed without increased thermal energy.
The present invention provides an apparatus and methods of converting electrons into x-rays within an imaging tube. An imaging tube is provided including a cathode and an anode. The cathode includes an emission surface, which emits a plurality of electrons along an emission axis. The anode includes a body having a track on a peripheral section of the body. The plurality of electrons are directed to impinge on the track at an impingement angle approximately equal to or between 15xc2x0 and 25xc2x0 relative to the emission axis and are converted into x-rays. A method of generating x-rays within the imaging tube is also provided.
One of several advantages of the present invention is that it provides an apparatus for emitting x-rays from an imaging tube with increased speed due to the ability to rotate the anode at increased speeds over traditional rotating anodes speeds.
Another advantage of the present invention is that due to mechanical and thermal operation of the imaging tube the present invention minimizes heat generated within the imaging tube as well as providing cooling of the anode while operating at the increased rotational speeds.
Furthermore, the present invention provides a smaller size anode, thus, reducing space requirements of the anode and increasing versatility as to application use of the imaging tube.